Flow estimation using hall-effect sensors

ABSTRACT

Methods for estimating flow rate in a blood circulation assist system employ impeller eccentricity. A method includes magnetically rotating an impeller within a blood flow channel of a blood pump. The impeller is levitated within the blood flow channel transverse to the impeller axis of rotation. A rotational speed for the impeller is determined. At least one impeller transverse position parameter is determined. The at least one impeller transverse position parameter is based on at least one of (1) an amount of a bearing current that is used to levitate the impeller transverse to the impeller axis of rotation, and (2) a position of the impeller within the blood flow channel transverse to the impeller axis of rotation. A flow rate of blood pumped by the blood pump is estimated based on the impeller rotational speed and the at least one impeller transverse position parameter.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional Application No.62/194,608 filed Jul. 20, 2015, the entire contents of which areincorporated herein by reference.

BACKGROUND

Ventricular assist devices, known as VADs, often include an implantableblood pump and are used for both short-term (i.e., days, months) andlong-term applications (i.e., years or a lifetime) where a patient'sheart is incapable of providing adequate circulation, commonly referredto as heart failure or congestive heart failure. According to theAmerican Heart Association, more than five million Americans are livingwith heart failure, with about 670,000 new cases diagnosed every year.People with heart failure often have shortness of breath and fatigue.Years of living with blocked arteries and/or high blood pressure canleave a heart too weak to pump enough blood to the body. As symptomsworsen, advanced heart failure develops.

A patient suffering from heart failure may use a VAD while awaiting aheart transplant or as a long term destination therapy. A patient mayalso use a VAD while recovering from heart surgery. Thus, a VAD cansupplement a weak heart (i.e., partial support) or can effectivelyreplace the natural heart's function.

The flow rate of blood pumped by a VAD is an important parameter forboth control of the blood pump and for informing a health careprofessional regarding the level of circulatory support provided to thepatient by the VAD. Direct blood flow rate measurement, however, may beundesirable with respect to additional components, such a flow ratesensor, that would be used to directly measure the flow rate of bloodpumped by the VAD. Such additional components may add to the complexityand size of the VAD, thereby potentially making the VAD more expensiveand occupy more space within the patient. Additionally, a flow ratesensor may increase the rate of thrombosis (blood clot formation) as aresult of the interface between the flow rate sensor and the blood flow.

In view of the challenges associated with direct blood flow ratemeasurement in a VAD, flow rate in a VAD may be estimated. For example,the blood flow rate in a VAD can be estimated based on the amount ofelectrical power consumed by the VAD. For some operational regimes of ablood pump, however, estimated flow rate based on electrical powerconsumed by the VAD may not be sufficiently accurate. As such, improvedapproaches for estimating blood flow rate in a VAD are desirable.

BRIEF SUMMARY

The following presents a simplified summary of some embodiments of theinvention in order to provide a basic understanding of the invention.This summary is not an extensive overview of the invention. It is notintended to identify key/critical elements of the invention or todelineate the scope of the invention. Its sole purpose is to presentsome embodiments of the invention in a simplified form as a prelude tothe more detailed description that is presented later.

Improved methods for estimating blood flow rate in a blood circulationassist system include determining an impeller position parameter. Incertain operating regimes of a blood pump, the impeller positionparameter and impeller rotational speed are used to estimate the bloodflow rate, thereby more accurately estimating the blood flow rate asopposed to estimating the blood flow rate based solely on impellerrotational rate and electrical power consumed by the blood pump.

Thus, in one aspect, a blood circulation assist system is provided thatestimates the flow rate of blood pumped based in part on impellerposition. The system includes a blood pump and a controller operativelycoupled with the blood pump. The blood pump includes an impellerdisposed within a blood flow channel of the blood pump and a motorstator operable to magnetically rotate the impeller. The impeller has animpeller axis of rotation around which the impeller is rotated. Themotor stator is operable to magnetically levitate the impeller withinthe blood flow channel transverse to the impeller axis of rotation. Thecontroller is configured to determine an impeller rotational speed forthe impeller, determine an amount of a drive current used to rotate theimpeller, and determine at least one impeller transverse positionparameter. The at least one impeller transverse position parameter isbased on at least one of (1) an amount of a bearing current that is usedto levitate the impeller transverse to the impeller axis of rotation,and (2) a position of the impeller within the blood flow channeltransverse to the impeller axis of rotation. The controller isconfigured to estimate a flow rate of blood pumped by the blood pumpbased on the impeller rotational speed and the drive current when thedrive current is below a first drive current threshold. The controlleris configured to estimate the flow rate based on the impeller rotationalspeed and the at least one impeller transverse position parameter whenthe drive current is above the first drive current threshold.

In some or all operating regimes of the blood pump, the controller canestimate the flow rate based on the impeller rotational speed, the drivecurrent, and the at least one impeller position parameter. For example,the controller can be configured to estimate the flow rate based on theimpeller rotational speed, the drive current, and the at least oneimpeller transverse position parameter when the drive current is betweena second drive current threshold and a third drive current threshold.

Any suitable approach can be used to determine the first drive currentthreshold. In many embodiments, the first drive current threshold variesbased on the impeller rotational speed. In many embodiments, the firstdrive current threshold is based on characteristics of variation in theamount of bearing current used to levitate the impeller transverse tothe impeller axis of rotation in response to variation in the impellerrotational speed. For example, the first drive current threshold can beselected such that the amount of bearing current used to levitate theimpeller transverse to the impeller axis increases in response to adecrease in the impeller rotational speed for drive currents above thefirst drive current threshold.

In many embodiments, the impeller impels the blood centrifugally and theblood pumped by the blood pump is output in a direction transverse tothe impeller axis of rotation. In such embodiments, the non-symmetricnature of the blood flow output induces eccentricity in the transverseposition of the impeller that varies with respect to the flow rate ofthe blood pumped. The at least one impeller transverse positionparameter can be indicative of an amount of eccentricity of the impellerwithin the blood flow channel.

Any suitable approach can be used to determine the at least one impellertransverse position parameter. In many embodiments, the system caninclude at least one sensor generating output indicative of the positionof the impeller within the blood flow channel transverse to the impelleraxis of rotation. For example, the at least one sensor can include aplurality of hall sensors generating output indicative of magnetic fluxlevels of the motor stator that are indicative of the position of theimpeller within the blood flow channel transverse to the impeller axisof rotation.

In many embodiments, the controller operates the blood pump tosubstantially minimize power consumption. For example, the controllercan be configured to control the amount of a bearing current that isused to levitate the impeller transverse to the impeller axis ofrotation so as to substantially minimize power consumption of the bloodpump. In many embodiments, the controller is configured to controleccentricity of the impeller within the blood flow channel so as tosubstantially minimize power consumption of the blood pump. In manyembodiments, the flow rate is estimated based on a target or measuredeccentricity of the impeller within the blood flow channel when thedrive current is above the first drive current threshold.

In many embodiments, the controller is configured to estimate a pressuredifferential across the impeller based on the at least one impellertransverse position parameter. For example, the pressure differentialcan be a function an off-center position for the impeller to minimizebearing current.

In another aspect, a method is provided for estimating blood flow ratein a blood circulation assist system. The method includes magneticallyrotating an impeller around an impeller axis of rotation within a bloodflow channel of a blood pump. The impeller is magnetically levitatedwithin the blood flow channel transverse to the impeller axis ofrotation. A controller operatively coupled with the blood pumpdetermines an impeller rotational speed for the impeller. The controllerdetermines at least one impeller transverse position parameter. The atleast one impeller transverse position parameter is based on at leastone of (1) an amount of a bearing current that is used to levitate theimpeller transverse to the impeller axis of rotation, and (2) a targetor measured position of the impeller within the blood flow channeltransverse to the impeller axis of rotation. The controller estimates aflow rate of blood pumped by the blood pump based on the impellerrotational speed and the at least one impeller transverse positionparameter.

In many embodiments, the method further includes estimating flow ratefor some operating regimes based on a drive current used to rotate theimpeller. For example, the controller can determine an amount of a drivecurrent used to rotate the impeller. The controller can estimate a flowrate of blood pumped by the blood pump based on the impeller rotationalspeed and the drive current. In many embodiments, the flow rate isestimated: (1) based on the impeller rotational speed and the drivecurrent when the drive current is below a first drive current threshold,and (2) based on the impeller rotational speed and the at least oneimpeller transverse position parameter when the drive current is abovethe first drive threshold.

The first drive current can be determined using any suitable approach.For example, the method can include selecting the first drive currentsuch that the amount of bearing current used to levitate the impellertransverse to the impeller axis increases in response to a decrease inthe impeller rotational speed for drive currents above the first drivecurrent threshold.

In some or all operating regimes of the blood pump, the method canestimate the flow rate based on the impeller rotational speed, the drivecurrent, and the at least one impeller transverse position parameter.For example, the method can include determining, with the controller, anamount of a drive current used to rotate the impeller. The controllercan estimate the flow rate based on the impeller rotational speed, thedrive current, and the at least one impeller transverse positionparameter when the drive current is between a second drive currentthreshold and a third drive current threshold.

In many embodiments of the method, the blood pump is controlled tosubstantially minimize power consumption of the blood pump. For example,the method can include controlling the amount of the bearing currentused to levitate the impeller transverse to the impeller axis ofrotation so as to substantially minimize power consumption of the bloodpump.

In many embodiments of the method, the blood pump is configured andoperated so that the transverse position of the impeller within theblood flow channel varies as a function of flow rate of the blood pumpfor at least a range of blood flow rates. For example, in manyembodiments of the method, the blood pump is configured such that theimpeller impels the blood centrifugally and the blood pumped by theblood pump is output in a direction transverse to the impeller axis ofrotation.

Any suitable approach can be used to determine the at least one impellertransverse position parameter used to estimate flow rate. For example,the method can include processing, with the controller, output from aplurality of hall sensors indicative of magnetic flux levels used tolevitate the impeller within the blood flow channel to determineeccentricity of the impeller within the blood flow channel. In manyembodiments, the at least one impeller transverse position parameter isindicative of the determined eccentricity or a target eccentricity.

In many embodiments, the method includes estimating, with thecontroller, a pressure differential across the impeller based on the atleast one impeller transverse position parameter. For example, thecontroller can estimate the pressure differential across the impellerbased on an off-center position for the impeller to minimize bearingcurrent.

For a fuller understanding of the nature and advantages of the presentinvention, reference should be made to the ensuing detailed descriptionand accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is an illustration of a mechanical circulatory support systemimplanted in a patient's body, in accordance with many embodiments.

FIG. 2 is an exploded view of certain components of the circulatorysupport system of FIG. 1.

FIG. 3 is an illustration of a blood pump in an operational positionimplanted in a patient's body.

FIG. 4 is a cross-sectional view of the blood pump of FIG. 3.

FIG. 5 is a partial cut-away perspective view of a stator of a bloodpump.

FIG. 6 is an illustration of an embodiment of a Hall Sensor assembly forthe blood pump of FIG. 3.

FIG. 7 is a schematic diagram of a control system architecture of themechanical support system of FIG. 1.

FIG. 8 graphically illustrates deviations between flow rates estimatedbased on power consumption and measured flow rates for an example bloodpump.

FIG. 9 shows a cross-sectional view of a centrifugal blood pump, inaccordance with many embodiments.

FIG. 10 schematically illustrates impeller eccentricity, in accordancewith many embodiments.

FIG. 11 shows a simplified schematic view of an impeller levitated via amotor stator, in accordance with many embodiments.

FIG. 12 is a plot showing observed correlations between measured flowrate and pump parameters including torque and impeller eccentricityvalues.

FIG. 13 is a simplified schematic illustration of a control architecturefor generating current applied to levitation coils of a blood pump totransversely levitate an impeller, in accordance with many embodiments.

FIG. 14 is a plot showing an observed correlation between measured flowand impeller drive current for an example blood pump operated at 3000rpm.

FIG. 15 is a plot showing an observed correlation between measured flowand target transverse eccentricity of the impeller of the example bloodpump of FIG. 14 operated at 3000 rpm.

FIG. 16 is a plot showing an observed correlation between measured flowand impeller drive current for the example blood pump of FIG. 14operated at 9000 rpm.

FIG. 17 is a plot showing an observed correlation between measured flowand target transverse eccentricity of the impeller of the example bloodpump of FIG. 14 operated at 9000 rpm.

FIG. 18 is a plot showing bearing current variations during pulsatilemode operation of the example blood pump of FIG. 14 operated at anominal 7000 rpm and 10 L/min.

FIG. 19 is a plot showing bearing current variations during pulsatilemode operation of the example blood pump of FIG. 14 operated at anominal 7000 rpm and 7 L/min.

FIG. 20 is a plot showing bearing current variations during pulsatilemode operation of the example blood pump of FIG. 14 operated at anominal 7000 rpm and 3 L/min.

FIG. 21 is a plot showing an observed correlation between measured flowand impeller drive current for the example blood pump of FIG. 14operated at 7000 rpm.

FIG. 22 is a plot showing an observed correlation between measured flowand target transverse eccentricity of the impeller of the example bloodpump of FIG. 14 operated at 7000 rpm.

FIG. 23 graphically illustrates deviations between flow rates estimatedbased on power consumption and target impeller eccentricity and measuredflow rates for the example blood pump of FIG. 14, in accordance withmany embodiments.

DETAILED DESCRIPTION

In the following description, various embodiments of the presentinvention will be described. For purposes of explanation, specificconfigurations and details are set forth in order to provide a thoroughunderstanding of the embodiments. However, it will also be apparent toone skilled in the art that the present invention may be practicedwithout the specific details. Furthermore, well-known features may beomitted or simplified in order not to obscure the embodiment beingdescribed.

Referring now to the drawings, in which like reference numeralsrepresent like parts throughout the several views, FIG. 1 is anillustration of a mechanical circulatory support system 10 implanted ina patient's body 12. The mechanical circulatory support system 10includes an implantable blood pump assembly 14, a ventricular cuff 16,an outflow cannula 18, an external system controller 20, and powersources 22. The implantable blood pump assembly 14 can include a VADthat is attached to an apex of the left ventricle, as illustrated, orthe right ventricle, or both ventricles of the heart 24. The VAD caninclude a centrifugal pump (as shown) that is capable of pumping theentire output delivered to the left ventricle from the pulmonarycirculation (i.e., up to 10 liters per minute). Related blood pumpsapplicable to the present invention are described in greater detailbelow and in U.S. Pat. Nos. 5,695,471, 6,071,093, 6,116,862, 6,186,665,6,234,772, 6,264,635, 6,688,861, 7,699,586, 7,976,271, 7,997,854,8,007,254, 8,152,493, 8,652,024, and 8,668,473 and U.S. PatentPublication Nos. 2007/0078293, 2008/0021394, 2009/0203957, 2012/0046514,2012/0095281, 2013/0096364, 2013/0170970, 2013/0121821, and2013/0225909, all of which are incorporated herein by reference for allpurposes in their entirety. With reference to FIGS. 1 and 2, the bloodpump assembly 14 can be attached to the heart 24 via the ventricularcuff 16, which can be sewn to the heart 24 and coupled to the blood pump14. The other end of the blood pump 14 connects to the ascending aortavia the outflow cannula 18 so that the VAD effectively diverts bloodfrom the weakened ventricle and propels it to the aorta for circulationthrough the rest of the patient's vascular system.

FIG. 1 illustrates the mechanical circulatory support system 10 duringbattery 22 powered operation. A driveline 26 that exits through thepatient's abdomen 28 connects the implanted blood pump assembly 14 tothe external system controller 20, which monitors system 10 operation.Related controller systems applicable to the present invention aredescribed in greater detail below and in U.S. Pat. Nos. 5,888,242,6,991,595, 8,323,174, 8,449,444, 8,506,471, 8,597,350, and 8,657,733, EP1812094, and U.S. Patent Publication Nos. 2005/0071001 and 2013/0314047,all of which are incorporated herein by reference for all purposes intheir entirety. The system 10 can be powered by either one, two, or morebatteries 22. It will be appreciated that although the system controller20 and power source 22 are illustrated outside/external to the patientbody, the driveline 26, the system controller 20 and/or the power source22 can be partially or fully implantable within the patient, as separatecomponents or integrated with the blood pump assembly 14. Examples ofsuch modifications are further described in U.S. Pat. No. 8,562,508 andU.S. Patent Publication No. 2013/0127253, all of which are incorporatedherein by reference for all purposes in their entirety.

With reference to FIGS. 3 to 5, a left ventricular assist blood pumpassembly 100 having a circular shaped housing 110 is implanted in apatient's body with a first face 111 of the housing 110 positionedagainst the patient's heart H and a second face 113 of the housing 110facing away from the heart H. The first face 111 of the housing 110includes an inlet cannula 112 extending into the left ventricle LV ofthe heart H. The second face 113 of the housing 110 has a chamfered edge114 to avoid irritating other tissue that may come into contact with theblood pump assembly 100, such as the patient's diaphragm. To constructthe illustrated shape of the puck-shaped housing 110 in a compact form,a stator 120 and electronics 130 of the pump assembly 100 are positionedon the inflow side of the housing toward first face 111, and a rotor 140of the pump assembly 100 is positioned along the second face 113. Thispositioning of the stator 120, electronics 130, and rotor 140 permitsthe edge 114 to be chamfered along the contour of the rotor 140, asillustrated in at least FIGS. 2-4, for example.

Referring to FIG. 4, the blood pump assembly 100 includes a dividingwall 115 within the housing 110 defining a blood flow conduit 103. Theblood flow conduit 103 extends from an inlet opening 101 of the inletcannula 112 through the stator 120 to an outlet opening 105 defined bythe housing 110. The rotor 140 is positioned within the blood flowconduit 103. The stator 120 is disposed circumferentially about a firstportion 140 a of the rotor 140, for example about a permanent magnet141. The stator 120 is also positioned relative to the rotor 140 suchthat, in use, blood flows within the blood flow conduit 103 through thestator 120 before reaching the rotor 140. The permanent magnet 141 has apermanent magnetic north pole N and a permanent magnetic south pole Sfor combined active and passive magnetic levitation of the rotor 140 andfor rotation of the rotor 140. The rotor 140 also has a second portion140 b that includes impeller blades 143. The impeller blades 143 arelocated within a volute 107 of the blood flow conduit such that theimpeller blades 143 are located proximate to the second face 113 of thehousing 110.

The puck-shaped housing 110 further includes a peripheral wall 116 thatextends between the first face 111 and a removable cap 118. Asillustrated, the peripheral wall 116 is formed as a hollow circularcylinder having a width W between opposing portions of the peripheralwall 116. The housing 110 also has a thickness T between the first face111 and the second face 113 that is less than the width W. The thicknessT is from about 0.5 inches to about 1.5 inches, and the width W is fromabout 1 inch to about 4 inches. For example, the width W can beapproximately 2 inches, and the thickness T can be approximately 1 inch.

The peripheral wall 116 encloses an internal compartment 117 thatsurrounds the dividing wall 115 and the blood flow conduit 103, with thestator 120 and the electronics 130 disposed in the internal compartment117 about the dividing wall 115. The removable cap 118 includes thesecond face 113, the chamfered edge 114, and defines the outlet opening105. The cap 118 can be threadedly engaged with the peripheral wall 116to seal the cap 118 in engagement with the peripheral wall 116. The cap118 includes an inner surface 118 a of the cap 118 that defines thevolute 107 that is in fluid communication with the outlet opening 105.

Within the internal compartment 117, the electronics 130 are positionedadjacent to the first face 111 and the stator 120 is positioned adjacentto the electronics 130 on an opposite side of the electronics 130 fromthe first face 111. The electronics 130 include circuit boards 131 andvarious components carried on the circuit boards 131 to control theoperation of the pump 100 (e.g., magnetic levitation and/or drive of therotor) by controlling the electrical supply to the stator 120. Thehousing 110 is configured to receive the circuit boards 131 within theinternal compartment 117 generally parallel to the first face 111 forefficient use of the space within the internal compartment 117. Thecircuit boards also extend radially-inward towards the dividing wall 115and radially-outward towards the peripheral wall 116. For example, theinternal compartment 117 is generally sized no larger than necessary toaccommodate the circuit boards 131, and space for heat dissipation,material expansion, potting materials, and/or other elements used ininstalling the circuit boards 131. Thus, the external shape of thehousing 110 proximate the first face 111 generally fits the shape of thecircuits boards 131 closely to provide external dimensions that are notmuch greater than the dimensions of the circuit boards 131.

With continued reference to FIGS. 4 and 5, the stator 120 includes aback iron 121 and pole pieces 123 a-123 f arranged at intervals aroundthe dividing wall 115. The back iron 121 extends around the dividingwall 115 and is formed as a generally flat disc of a ferromagneticmaterial, such as steel, in order to conduct magnetic flux. The backiron 121 is arranged beside the control electronics 130 and provides abase for the pole pieces 123 a-123 f.

Each of the pole piece 123 a-123 f is L-shaped and has a drive coil 125for generating an electromagnetic field to rotate the rotor 140. Forexample, the pole piece 123 a has a first leg 124 a that contacts theback iron 121 and extends from the back iron 121 towards the second face113. The pole piece 123 a can also have a second leg 124 b that extendsfrom the first leg 124 a through an opening of a circuit board 131towards the dividing wall 115 proximate the location of the permanentmagnet 141 of the rotor 140. In an aspect, each of the second legs 124 bof the pole pieces 123 a-123 f is sticking through an opening of thecircuit board 131. In an aspect, each of the first legs 124 a of thepole pieces 123 a-123 f is sticking through an opening of the circuitboard 131. In an aspect, the openings of the circuit board are enclosingthe first legs 124 a of the pole pieces 123 a-123 f.

In a general aspect, the implantable blood pump 100 can include one ormore Hall sensors that may provide an output voltage, which is directlyproportional to a strength of a magnetic field that is located inbetween at least one of the pole pieces 123 a-123 f and the permanentmagnet 141, and the output voltage may provide feedback to the controlelectronics 130 of the pump 100 to determine if the rotor 140 and/or thepermanent magnet 141 is not at its intended position for the operationof the pump 100. For example, a position of the rotor 140 and/or thepermanent magnet 141 can be adjusted, e.g., the rotor 140 or thepermanent magnet 141 may be pushed or pulled towards a center of theblood flow conduit 103 or towards a center of the stator 120.

Each of the pole pieces 123 a-123 f also has a levitation coil 127 forgenerating an electromagnetic field to control the radial position ofthe rotor 140. Each of the drive coils 125 and the levitation coils 127includes multiple windings of a conductor around the pole pieces 123a-123 f. Particularly, each of the drive coils 125 is wound around twoadjacent ones of the pole pieces 123, such as pole pieces 123 d and 123e, and each levitation coil 127 is wound around a single pole piece. Thedrive coils 125 and the levitation coils 127 are wound around the firstlegs of the pole pieces 123, and magnetic flux generated by passingelectrical current though the coils 125 and 127 during use is conductedthrough the first legs and the second legs of the pole pieces 123 andthe back iron 121. The drive coils 125 and the levitation coils 127 ofthe stator 120 are arranged in opposing pairs and are controlled todrive the rotor and to radially levitate the rotor 140 by generatingelectromagnetic fields that interact with the permanent magnetic poles Sand N of the permanent magnet 141. Because the stator 120 includes boththe drive coils 125 and the levitation coils 127, only a single statoris needed to levitate the rotor 140 using only passive and activemagnetic forces. The permanent magnet 141 in this configuration has onlyone magnetic moment and is formed from a monolithic permanent magneticbody 141. For example, the stator 120 can be controlled as discussed inU.S. Pat. No. 6,351,048, the entire contents of which are incorporatedherein by reference for all purposes. The control electronics 130 andthe stator 120 receive electrical power from a remote power supply via acable 119 (FIG. 3). Further related patents, namely U.S. Pat. Nos.5,708,346, 6,053,705, 6,100,618, 6,222,290, 6,249,067, 6,278,251,6,351,048, 6,355,998, 6,634,224, 6,879,074, and 7,112,903, all of whichare incorporated herein by reference for all purposes in their entirety.

The rotor 140 is arranged within the housing 110 such that its permanentmagnet 141 is located upstream of impeller blades in a location closerto the inlet opening 101. The permanent magnet 141 is received withinthe blood flow conduit 103 proximate the second legs 124 b of the polepieces 123 to provide the passive axial centering force thoughinteraction of the permanent magnet 141 and ferromagnetic material ofthe pole pieces 123. The permanent magnet 141 of the rotor 140 and thedividing wall 115 form a gap 108 between the permanent magnet 141 andthe dividing wall 115 when the rotor 140 is centered within the dividingwall 115. The gap 108 may be from about 0.2 millimeters to about 2millimeters. For example, the gap 108 can be approximately 1 millimeter.The north permanent magnetic pole N and the south permanent magneticpole S of the permanent magnet 141 provide a permanent magneticattractive force between the rotor 140 and the stator 120 that acts as apassive axial centering force that tends to maintain the rotor 140generally centered within the stator 120 and tends to resist the rotor140 from moving towards the first face 111 or towards the second face113. When the gap 108 is smaller, the magnetic attractive force betweenthe permanent magnet 141 and the stator 120 is greater, and the gap 108is sized to allow the permanent magnet 141 to provide the passivemagnetic axial centering force having a magnitude that is adequate tolimit the rotor 140 from contacting the dividing wall 115 or the innersurface 118 a of the cap 118. The rotor 140 also includes a shroud 145that covers the ends of the impeller blades 143 facing the second face113 that assists in directing blood flow into the volute 107. The shroud145 and the inner surface 118 a of the cap 118 form a gap 109 betweenthe shroud 145 and the inner surface 118 a when the rotor 140 islevitated by the stator 120. The gap 109 is from about 0.2 millimetersto about 2 millimeters. For example, the gap 109 is approximately 1millimeter.

As blood flows through the blood flow conduit 103, blood flows through acentral aperture 141 a formed through the permanent magnet 141. Bloodalso flows through the gap 108 between the rotor 140 and the dividingwall 115 and through the gap 109 between the shroud 145 and the innersurface 108 a of the cap 118. The gaps 108 and 109 are large enough toallow adequate blood flow to limit clot formation that may occur if theblood is allowed to become stagnant. The gaps 108 and 109 are also largeenough to limit pressure forces on the blood cells such that the bloodis not damaged when flowing through the pump 100. As a result of thesize of the gaps 108 and 109 limiting pressure forces on the bloodcells, the gaps 108 and 109 are too large to provide a meaningfulhydrodynamic suspension effect. That is to say, the blood does not actas a bearing within the gaps 108 and 109, and the rotor is onlymagnetically-levitated. In various embodiments, the gaps 108 and 109 aresized and dimensioned so the blood flowing through the gaps forms a filmthat provides a hydrodynamic suspension effect. In this manner, therotor can be suspended by magnetic forces, hydrodynamic forces, or both.

Because the rotor 140 is radially suspended by active control of thelevitation coils 127 as discussed above, and because the rotor 140 isaxially suspended by passive interaction of the permanent magnet 141 andthe stator 120, no rotor levitation components are needed proximate thesecond face 113. The incorporation of all the components for rotorlevitation in the stator 120 (i.e., the levitation coils 127 and thepole pieces 123) allows the cap 118 to be contoured to the shape of theimpeller blades 143 and the volute 107. Additionally, incorporation ofall the rotor levitation components in the stator 120 eliminates theneed for electrical connectors extending from the compartment 117 to thecap 118, which allows the cap to be easily installed and/or removed andeliminates potential sources of pump failure.

In use, the drive coils 125 of the stator 120 generates electromagneticfields through the pole pieces 123 that selectively attract and repelthe magnetic north pole N and the magnetic south pole S of the rotor 140to cause the rotor 140 to rotate within stator 120. For example, the oneor more Hall sensors may sense a current position of the rotor 140and/or the permanent magnet 141, wherein the output voltage of the oneor more Hall sensors may be used to selectively attract and repel themagnetic north pole N and the magnetic south pole S of the rotor 140 tocause the rotor 140 to rotate within stator 120. As the rotor 140rotates, the impeller blades 143 force blood into the volute 107 suchthat blood is forced out of the outlet opening 105. Additionally, therotor draws blood into pump 100 through the inlet opening 101. As bloodis drawn into the blood pump by rotation of the impeller blades 143 ofthe rotor 140, the blood flows through the inlet opening 101 and flowsthrough the control electronics 130 and the stator 120 toward the rotor140. Blood flows through the aperture 141 a of the permanent magnet 141and between the impeller blades 143, the shroud 145, and the permanentmagnet 141, and into the volute 107. Blood also flows around the rotor140, through the gap 108 and through the gap 109 between the shroud 145and the inner surface 118 a of the cap 118. The blood exits the volute107 through the outlet opening 105, which may be coupled to an outflowcannula.

FIG. 6 shows a Hall Sensor assembly 200 for the blood pump assembly 14,in accordance with many embodiments. The Hall Sensor assembly 200includes a printed circuit board (PCB) 202 and individual Hall Effectsensors 208 supported by the printed circuit board 202. Eightaxi-symmetric Hall Effect sensors 208 are placed in a rigid, plasticmechanical carrier 210 and the PCB 202 is placed onto the mechanicalcarrier 210. The mechanical carrier 210 uses guide rails 212 to locateelectrically neutral rigid PCB portions 214 attached to the top edges ofthe Hall Effect sensors 208 and to locate the PCB 202.

The Hall Effect sensors 208 are configured to transduce a position ofthe rotor 140 of the pump 100. In the illustrated embodiment, the HallEffect sensors 208 are supported so as to be standing orthogonallyrelative to the PCB 202 and a longest edge of each of the Hall Effectsensors 208 is aligned to possess an orthogonal component with respectto the surface of the PCB 202. Each of the Hall Effect sensors 208generate an output voltage, which is directly proportional to a strengthof a magnetic field that is located in between at least one of the polepieces 123 a-123 f and the permanent magnet 141. The voltage output byeach of the Hall Effect sensors 208 is received by the controlelectronics 130, which processes the sensor output voltages to determinethe position and orientation of the rotor 140. The determined positionand orientation of the rotor 140 is used to determine if the rotor 140is not at its intended position for the operation of the pump 100. Forexample, a position of the rotor 140 and/or the permanent magnet 141 maybe adjusted, for example, the rotor 140 or the permanent magnet 141 maybe pushed or pulled towards a center of the blood flow conduit 103 ortowards a center of the stator 120. The determined position of the rotor140 can also be used to determine rotor eccentricity or a target rotoreccentricity, which can be used as described herein to estimate flowrate of blood pumped by the blood pump assembly 100.

FIG. 7 is a schematic diagram of a control system architecture of themechanical support system of FIG. 1. The driveline 26 couples theimplanted blood pump assembly 100 to the external system controller 20,which monitors system operation via various software applications. Theblood pump assembly 100 itself also includes several softwareapplications that are executable by the on board electronics 130 (e.g.,processors) for various functions, such as to control radial levitationand/or drive of the rotor of the pump assembly 100 during operation. Theexternal system controller 20 can in turn be coupled to batteries 22 ora power module 30 that connect to an AC electrical outlet. The externalsystem controller 20 can also include an emergency backup battery (EBB)to power the system (e.g., when the batteries 22 are depleted) and amembrane overlay, including Bluetooth capabilities for wireless datacommunication. An external computer having a system monitor 32 that isconfigurable by an operator, such as clinician or patient, may furtherbe coupled to the circulatory support system for configuring theexternal system controller 20, implanted blood pump assembly 100, and/orpatient specific parameters, updating software on the external systemcontroller 20 and/or implanted blood pump assembly 100, monitoringsystem operation, and/or as a conduit for system inputs or outputs.

Impeller Position Based Flow Estimation

FIG. 8 graphically illustrates example deviations between measured flowrate for an example blood pump and flow rates (220-3000, 220-4000,220-5000, 220-6000, 220-7000, 220-8000, 220-9000) estimated for theexample blood pump based on power consumption and rotor rotation ratefor a number of different impeller rotation rates. Ideally, theestimated flow rate would correspond to an exact estimated flow 222 thatis equal to the measured flow rate. For a range of impeller rotationrates, however, there is a range of measured flow rates in which, for aparticular impeller rotation rate and a particular blood pump, the samepower consumption magnitude is used to produce two different actualmeasure flow rates. For example, for the estimated flow rate curve220-9000 for an impeller rotation rate of 9000 rpm, the same estimatedflow rate of 8.0 L/min corresponds to two different actual measured flowrates of about 7.8 L/min and 10.6 L/min. Moreover, for all actualmeasured flow rates above 9.3 L/min for an impeller rotation rate of9000 rpm, the actual power consumption drops with increasing flow ratethereby resulting in increasing magnitude of error between the estimateflow rate 220-9000 and the exact estimated flow 222. Such a doubledvalue characteristic can also be seen in the data displayed in FIG. 8for impeller rotation rates of 4000 rpm to 9000 rpm. Accordingly,estimating flow rate based only on impeller rotation rate and powerconsumption can result in significant relative error for actual pumpflow rates at the high end of the actual flow rate range. In manyembodiments, one or more additional flow rate related parameters areemployed to produce more accurate flow rate estimates. As describedherein, in many embodiments, estimated flow rate for at least someranges of flow rate is based on an actual or target impeller transverseeccentricity.

FIG. 9 shows a cross-sectional view of the centrifugal blood pump 14, inaccordance with many embodiments. As described herein, the rotor 140 ismagnetically levitated in the blood flow channel 103 via magneticinteraction between the permanent magnet 141 and the motor stator 120.The blood flows through the center of the rotor 140 and is impelled intoan axially non-symmetric output chamber 224 from which the blood isoutput from the blood pump 14 via the outlet 105 shown in FIG. 4. As aresult of the transverse momentum imparted to the blood flow by therotor 140, the rotor is subjected to a transverse force that is reactedvia magnetic interaction between the permanent magnet 141 and the motorstator 120. While the motor stator 120 can be controlled so as to keepthe rotor 140 centered within the blood flow channel, the powerconsumption of the blood pump 14 can be reduced by allowing the positionof the rotor 140 to deviate from being centered in the blood flowchannel 103. Moreover, as described herein, minimizing the powerconsumption of the blood pump 14 over a range of operating conditions(e.g., impeller rotation rate, flow rate) results in target impellereccentricity that varies as a function of flow rate over at least arange of operating conditions so as to enable use of the target impellereccentricity as a parameter from which to estimate flow rate. FIG. 10schematically illustrates impeller eccentricity that occurs in thepresence of transverse force being applied to the rotor 140 andminimization of the power consumption of the blood pump 14.

FIG. 11 shows a simplified cross-sectional schematic view of the rotor140 levitated via the motor stator 120. The output from the Hall sensors208 is processed to determine the transverse position of the rotor 140in both X and Y directions. The displacement of the rotor 140 from acentered reference position in each of the X and Y directions can becombined to generate a total vector sum eccentricity of the rotor 140from the centered reference position.

FIG. 12 is a plot showing observed correlations between measured flowrate and pump parameters including torque and impeller eccentricityvalues for an example pump operated at 9000 rpm. As can be seen, thetorque 226 increases with flow up to a flow rate of about 9.5 L/min andthen drops thereafter. The torque 226 is proportional to drive currentand is double valued in the shaded range of flow rate greater than 7.5L/min. Accordingly, estimation of pump flow rate based on the torque 226for flow rates greater than about 9.5 L/min may produce increasinglymore relative error between the estimated flow rate and the actual flowrate. In contrast, the total rotor eccentricity 228, which is the vectorsum of the X-direction rotor eccentricity 230 and the Y-direction rotoreccentricity 232, is single valued in the shaded range of flow rategreater than 7.5 L/min, and can therefore be employed to estimate flowrate at least in some or all of the shaded flow rate range of greaterthan 7.5 L/min.

As described herein, the blood pump 14 magnetically levitates androtates the rotor 140. Driving current is applied to the drive coils125. Current for levitating the rotor is applied to the levitation coils127. FIG. 13 is a simplified schematic illustration of a bearing currentcontroller 250 for generating current applied to the levitation coils127 of the blood pump 14 to transversely levitate the rotor 140, inaccordance with many embodiments. The bearing current controller 250includes a magnetic center proportional-integral-derivative (PID)controller 252, a position PID controller 254, and a current PIDcontroller 256. The current generated is applied to the levitation coils127 to controllably levitate the rotor 140. The resulting position ofthe rotor 140 (signal 258) is determined from the output of the Hallsensors 208. The bearing current controller 250 employs a three-levelcascaded PhD control method. The ultimate feedback signal is the bearingcurrent 260 therefore the bearing current controller 250 is configuredto minimize bearing current to reduce power consumption of the bloodpump 14.

At different flow rates, the resulting bearing current is differentreflecting the different forces on the rotor from the impelled blood. Inthe described embodiment, there are two bearing currents because thereare two separate bearing coils on the stator 120 (two for eachdirection). A Park transformation is applied to change the coordinatesfrom stator referenced directions (X and Y) to the rotor referenceddirections (d and q directions). Two separate bearing currentcontrollers 250 are used to control the current applied to thelevitation coils 127—one for each of the d and q directions. Thedirection d is aligned along the rotor N-S dimension. The direction q isperpendicular to the direction d. The directions d and q define a planeperpendicular to the direction of flow through the center of the rotor140.

The magnetic center PID controller 252 generates reference signals forthe position PID controller 254 defining a target off-center positionfor the rotor (in the d-q coordinate system) to minimize bearingcurrent. The position PID controller 254 generates reference signals forthe bearing current in the d-q coordinate system. The current PIDcontroller 254 calculates the bearing current in the d-q coordinatesystem and then applies an inverse Park transformation to generatecurrent output for application to the levitation coils 127 to controllevitation of the rotor 140 in each of the two separate directions (Xand Y).

Signals 262, 264, 266, 268, 260, 258 generated by the bearing currentcontroller 250 were studied for possible use in estimating flow rate.Because of the cascaded control structure employed, the signals 262,264, 266, 268, 260, 258 generated by the bearing current controller 250show similar trend of changes when flow rate is changed, although thetrend direction may be reversed because of the negative feedback signchange. Signals 266, 268, 260, 258 show high run-to-run variation andnoise-to-signal ratio is high due to the bearing current and centerposition feedback signals have significant disturbance induced from thefluid field. Signal 262 and signal 264 are more stable because thefeedback signal is the bearing current after a low pass filter. Signal264 is the target reference rotor center position, which is even morestable than signal 262, because the gain in the magnetic center PIDcontroller 252 is zero. The magnetic center PID controller 252 imposes20 dB attenuation from DC up to a frequency. Accordingly, the targetreference rotor center position signal 264, which defines the targetoff-center position for the rotor (in the d-q coordinate system) tominimize bearing current, provides a suitable signal that can beprocessed to estimate flow rate of the blood pump. For example, each ofthe target reference rotor center signals 264 from the two bearingcurrent controllers 250 (one for the X-direction levitation current andone for the Y-direction levitation current) can be combined to calculatea target reference rotor center value corresponding to a total targeteccentric distance of the target off-center position for the rotor fromthe center of the blood flow channel of the blood pump.

FIG. 14 is a plot showing an observed correlation between measured flowand impeller drive current for an example blood pump operated at 3000rpm. As illustrated in FIG. 14, for measured flow rates from 1 L/min to2.5 L/min, the corresponding drive current used to rotate the impellershows a substantially linear increase from about 288 counts to about 318counts. Above 2.5 L/min, the increase in drive current with increasedmeasured flow rate diminishes down to no significant increase in drivecurrent for measured flow rate between 3.5 L/min and 4.0 L/min.Accordingly, estimated flow rate based only on drive current forimpeller rotation rate of 3000 rpm may deviate increasingly from actualflow rate for flow rates above 2.5 L/min.

FIG. 15 is a plot showing an observed correlation between measured flowand the target reference center value of the example blood pump of FIG.14 operated at 3000 rpm. As illustrated in FIG. 15, for measured flowrates from 1 L/min to 3.5 L/min, the corresponding target referencecenter value shows a substantially linear increase. Therefore, thetarget reference center value can be used to increase the accuracy ofestimated flow rate, at least in the 2.5 L/min to 3.5 L/min range forthe example blood pump of FIG. 14 operated at 3000 rpm.

FIG. 16 is a plot showing an observed correlation between measured flowand impeller drive current for the example blood pump of FIG. 14operated at 9000 rpm. As illustrated in FIG. 16, for measured flow ratesfrom about 1.5 L/min to about 10.0 L/min, the corresponding drivecurrent used to rotate the impeller shows a substantially linearincrease from about 1120 counts to about 1750 counts. From about 10.0L/min to about 11.0 L/min, the corresponding drive current used torotate the impeller shows no significant change. Above 11.0 L/min, thecorresponding drive current used to rotate the impeller drops down toabout 1600 counts at about 13.0 L/min flow rate. Accordingly, estimatedflow rate based only on drive current for impeller rotation rate of 9000rpm may deviate increasingly from actual flow rate for flow rates above10.0 L/min.

FIG. 17 is a plot showing an observed correlation between measured flowand the target reference center value of the example blood pump of FIG.14 operated at 9000 rpm. As illustrated in FIG. 17, for measured flowrates from 1 L/min to 8.0 L/min, the corresponding target referencecenter value shows a substantially linear decrease. From 8.0 L/min to9.0 L/min flow rate, the corresponding target reference center valueshows no significant change. Above 9.0 L/min flow rate, thecorresponding target reference center value exhibits a substantiallylinear increase with increasing flow rate. Therefore, the targetreference center value can be used to increase the accuracy of estimatedflow rate, at least in the 9.0 L/min and higher range for the exampleblood pump of FIG. 14 operated at 9000 rpm.

An interesting observation from FIG. 15 and FIG. 17 is that the targetreference center value is correlated with flow rate. Although the targetreference center value also has bell curve shape meaning it also hasdouble value problem for use in estimating flow rate, its double valuerange is different from the double value range of the driving current.Accordingly, one method that can be used to overcome the estimationerror arising from the double valued nature of the driving current andflow rate correlation is to combine the driving current with the targetreference center to predict the flow rate, for example, switching fromdriving current to the target reference center when driving current isin the double value range.

There are at least two approaches for selecting when to switch betweenestimating flow rate based on driving current and estimating flow ratebased on target reference center. One method selects a single flow ratevalue to switch for each particular rotor rotational rate. For example,when driving current is higher than 1600 counts in FIG. 16, drivingcurrent is in the double value range, so the flow can be estimated basedon target reference center instead of drive current when the drivingcurrent is higher than 1600 counts.

A second method for determining what flow rate to switch betweenestimating flow rate based on driving current and estimating flow ratebased on target reference center is based on how the bearing currentvaries in response to a pulsatile variation in the rotation rate of therotor. The merit of the second method is that no calibration variable isinvolved in the second method's algorithm based switching, which maytherefore be more robust in implementation. During pulsatile modeoperation of the blood pump, the rotor rotational rate is periodicallyvaried to simulate natural blood pulse. On a periodic basis, the rotorrotational rate is temporarily reduced from the current nominalrotational rate, then temporarily increased from the reduced rate to arate greater than the current nominal rotation rate, and then reducedback down to the current nominal rotational rate. During each of thesepulsatile rotation rate variations, a transition is also observed inbearing current. When bearing current is decomposed into d-qcoordinates, it is found that the pulsatile transition of the bearingcurrent changes with flow rate. FIGS. 18 through 20 show the bearingcurrents filtered by the low pass filter and decomposed into d-qcoordinates for different flow rates for a rotor rotation rate of 7000rpm. The left subfigure is for the d axis, and right for the q axis. Thed axis bearing current transition goes up then down at 10 L/min, andgoes down then up at 3 L/min. The change in direction of the bearingcurrent transition indicates that 10 L/min flow rate and 3 L/min flowrate are on different sides of the target reference center bell curve.FIG. 21 is a plot showing an observed correlation between measured flowand impeller drive current for the example blood pump of FIG. 14operated at 7000 rpm. FIG. 22 is a plot showing an observed correlationbetween measured flow and target reference center of the impeller of theexample blood pump of FIG. 14 operated at 7000 rpm. As shown in FIG. 22,the target reference center value bottoms at 6 L/min. Therefore theslope of the target reference center curve with flow is reversed at 6L/min, which causes the difference in pulsatile transition of therelated rotor levitation current.

This pulsatile transition change in bearing current can be used to setan estimation parameter having either a value of 0.0 or 1.0 based onwhether the shape of the pulsatile transition change in bearing currentindicates that the current nominal flow rate is lower or higher than theflow rate at which the target reference center curve bottoms. Theestimation parameter can then be used to switch between estimating flowrate based on drive current when the flow rate is below the flow rate atwhich the target reference center curve bottoms and estimating flow ratebased on target reference center value when the flow rate is above theflow rate at which the target reference center bottoms. For example,FIGS. 21 and 22 show that for an example pump operated at 7000 rpm,driving current is a good estimator until about 7 L/min. Based on thepulsatile transition change in bearing current signature, the estimationparameter can be set to 1.0 at 7 L/min and higher so that the targetreference center value can be used to estimate flow rate for flow ratesof 7 L/min or higher.

Calibration can also be done to fit the target reference center signalwith the flow rate. To avoid switching noise when operating close to theswitching flow rate, a weighing method can be used to put less weight onthe target reference center signal when the pulsatile transition changein the bearing current indicates a flow rate corresponding toapproximately the bottom of the target reference signal, and graduallyadd more weight to reference center signal at higher flow rate. Theaccuracy of this weighing method is illustrated in FIG. 23, whichgraphically illustrates deviations between the resulting measured flowrate and actual flow rate for the example blood pump of FIG. 14.

Other suitable approaches for increasing the accuracy of flow rateestimation using parameters related to rotor levitation are alsopossible. For example, any suitable existing curve fitting techniquescan be used to estimate flow rate based on any suitable combination ofrotor rotation rate, drive current, and target reference center. Also,two or more different curve fits can be used to cover the entire rangeof flow rates. For example, one curve fit can be used to estimate flowrate at the low range of flow rates where flow rate is primarily afunction of driving current, a second curve fit can be used at the highrange of flow rates based on target reference center, and a third curvefit can be used at the mid-range of flow rates based on both drivingcurrent and target flow rate. Other bearing current related parameterscan also be used. For example, the bearing current can be controlled tokeep the rotor centered in the blood flow channel and the variation inthe bearing current, which will be greater if the rotor is keptcentered, can be used as another parameter in addition to rotor drivingcurrent to estimate flow rate.

Impeller Position Based Pump Pressure Differential Estimation

One of skill in the art would appreciate that the parameters related toimpeller position described herein (e.g., bearing current, off-centerposition for the impeller to minimize bearing current) can be used aloneor in combination to estimate pressure differential across the impeller(i.e., difference in pressure on the output side of the impeller topressure on the input side of the impeller) in addition to or instead ofestimating flow. For example, in many embodiments of the centrifugalblood pump 14, the impeller eccentricity for minimum bearing currentappears to be solely or mostly dependent upon the pressure differentialacross the impeller. As a result, the pressure differential across theimpeller can be estimated using a suitable function of the parametersrelated to impeller position described herein. Also, any suitableadditional operational parameter of the blood pump, such as pumprotational speed, impeller drive current, and/or estimated blood flowthrough the pump, can be used alone or in any suitable combination inaddition to the parameters related to impeller position described hereinto estimate the pressure differential across the impeller. Moreover, oneof skill would appreciate that the pressure differential across theimpeller can be derived from the flow rate of blood through the bloodpump and vice versa. The resulting estimated pressure differential canbe used in any suitable way, including as a parameter on which operationof the pump is based to produce desired pressure differential across thepump suitable for particular patient physiological conditions and/or todetect and react to adverse pump conditions.

Other variations are within the spirit of the present invention. Thus,while the invention is susceptible to various modifications andalternative constructions, certain illustrated embodiments thereof areshown in the drawings and have been described above in detail. It shouldbe understood, however, that there is no intention to limit theinvention to the specific form or forms disclosed, but on the contrary,the intention is to cover all modifications, alternative constructions,and equivalents falling within the spirit and scope of the invention, asdefined in the appended claims.

The use of the terms “a” and “an” and “the” and similar referents in thecontext of describing the invention (especially in the context of thefollowing claims) are to be construed to cover both the singular and theplural, unless otherwise indicated herein or clearly contradicted bycontext. The terms “comprising,” “having,” “including,” and “containing”are to be construed as open-ended terms (i.e., meaning “including, butnot limited to,”) unless otherwise noted. The term “connected” is to beconstrued as partly or wholly contained within, attached to, or joinedtogether, even if there is something intervening. Recitation of rangesof values herein are merely intended to serve as a shorthand method ofreferring individually to each separate value falling within the range,unless otherwise indicated herein, and each separate value isincorporated into the specification as if it were individually recitedherein. All methods described herein can be performed in any suitableorder unless otherwise indicated herein or otherwise clearlycontradicted by context. The use of any and all examples, or exemplarylanguage (e.g., “such as”) provided herein, is intended merely to betterilluminate embodiments of the invention and does not pose a limitationon the scope of the invention unless otherwise claimed. No language inthe specification should be construed as indicating any non-claimedelement as essential to the practice of the invention.

Preferred embodiments of this invention are described herein, includingthe best mode known to the inventors for carrying out the invention.Variations of those preferred embodiments may become apparent to thoseof ordinary skill in the art upon reading the foregoing description. Theinventors expect skilled artisans to employ such variations asappropriate, and the inventors intend for the invention to be practicedotherwise than as specifically described herein. Accordingly, thisinvention includes all modifications and equivalents of the subjectmatter recited in the claims appended hereto as permitted by applicablelaw. Moreover, any combination of the above-described elements in allpossible variations thereof is encompassed by the invention unlessotherwise indicated herein or otherwise clearly contradicted by context.

All references, including publications, patent applications, andpatents, cited herein are hereby incorporated by reference to the sameextent as if each reference were individually and specifically indicatedto be incorporated by reference and were set forth in its entiretyherein.

What is claimed is:
 1. A blood circulation assist system, comprising: ablood pump including an impeller disposed within a blood flow channel ofthe blood pump and a motor stator operable to magnetically rotate theimpeller, the impeller having an impeller axis of rotation around whichthe impeller is rotated, the motor stator being further operable tomagnetically levitate the impeller within the blood flow channeltransverse to the impeller axis of rotation; a controller operativelycoupled with the blood pump, the controller being configured to:determine an impeller rotational speed for the impeller; determine anamount of a drive current used to rotate the impeller; determine atleast one impeller transverse position parameter, the at least oneimpeller transverse position parameter being based on at least one of(1) an amount of a bearing current that is used to levitate the impellertransverse to the impeller axis of rotation, and (2) a position of theimpeller within the blood flow channel transverse to the impeller axisof rotation; estimate a flow rate of blood pumped by the blood pumpbased on the impeller rotational speed and the drive current when thedrive current is below a first drive current threshold; and estimate theflow rate based on the impeller rotational speed and the at least oneimpeller transverse position parameter when the drive current is abovethe first drive current threshold.
 2. The blood circulation assistsystem of claim 1, wherein the controller is configured to estimate theflow rate based on the impeller rotational speed, the drive current, andthe at least one impeller transverse position parameter when the drivecurrent is between a second drive current threshold and a third drivecurrent threshold.
 3. The blood circulation assist system of claim 1,wherein the first drive current threshold varies based on the impellerrotational speed.
 4. The blood circulation assist system of claim 1,wherein the first drive current threshold is based on characteristics ofvariation in the amount of bearing current used to levitate the impellertransverse to the impeller axis of rotation in response to variation inthe impeller rotational speed.
 5. The blood circulation assist system ofclaim 4, wherein the first drive current threshold is selected such thatthe amount of bearing current used to levitate the impeller transverseto the impeller axis increases in response to a decrease in the impellerrotational speed for drive currents above the first drive currentthreshold.
 6. The blood circulation assist system of claim 1, whereinthe impeller impels the blood centrifugally and the blood pumped by theblood pump is output in a direction transverse to the impeller axis ofrotation.
 7. The blood circulation assist system of claim 1, the atleast one impeller transverse position parameter is indicative of anamount of eccentricity of the impeller within the blood flow channel. 8.The blood circulation assist system of claim 1, further comprising atleast one sensor generating output indicative of the position of theimpeller within the blood flow channel transverse to the impeller axisof rotation.
 9. The blood circulation system of claim 8, wherein the atleast one sensor comprises a plurality of hall sensors generating outputindicative of magnetic flux levels of the motor stator.
 10. The bloodcirculation assist system of claim 1, wherein the controller isconfigured to control the amount of a bearing current that is used tolevitate the impeller transverse to the impeller axis of rotation so asto substantially minimize power consumption of the blood pump.
 11. Theblood circulation assist system of claim 1, wherein the controller isconfigured to control eccentricity of the impeller within the blood flowchannel so as to substantially minimize power consumption of the bloodpump.
 12. The blood circulation assist system of claim 11, wherein theflow rate is estimated based on eccentricity of the impeller within theblood flow channel when the drive current is above the first drivecurrent threshold.
 13. The blood circulation assist system of claim 1,wherein the controller is configured to estimate a pressure differentialacross the impeller based on the at least one impeller transverseposition parameter.
 14. A method for estimating blood flow rate in ablood circulation assist system, the method comprising: magneticallyrotating an impeller around an impeller axis of rotation within a bloodflow channel of a blood pump; magnetically levitating the impellerwithin the blood flow channel transverse to the impeller axis ofrotation; determining, with a controller operatively coupled with theblood pump, an impeller rotational speed for the impeller; determining,with the controller, at least one impeller transverse positionparameter, the at least one impeller transverse position parameter beingbased on at least one of (1) an amount of a bearing current that is usedto levitate the impeller transverse to the impeller axis of rotation,and (2) a position of the impeller within the blood flow channeltransverse to the impeller axis of rotation; and estimating, with thecontroller, a flow rate of blood pumped by the blood pump based on theimpeller rotational speed and the at least one impeller transverseposition parameter.
 15. The method of claim 14, further comprising:determining, with the controller, an amount of a drive current used torotate the impeller; and estimating, with the controller, a flow rate ofblood pumped by the blood pump based on the impeller rotational speedand the drive current.
 16. The method of claim 15, wherein the flow rateis estimated: based on the impeller rotational speed and the drivecurrent when the drive current is below a first drive current threshold;and based on the impeller rotational speed and the at least one impellertransverse position parameter when the drive current is above the firstdrive threshold.
 17. The method of claim 16, comprising selecting thefirst drive current threshold such that the amount of bearing currentused to levitate the impeller transverse to the impeller axis increasesin response to a decrease in the impeller rotational speed for drivecurrents above the first drive current threshold.
 18. The method ofclaim 14, further comprising: determining, with the controller, anamount of a drive current used to rotate the impeller; and estimating,with the controller, the flow rate based on the impeller rotationalspeed, the drive current, and the at least one impeller transverseposition parameter when the drive current is between a second drivecurrent threshold and a third drive current threshold.
 19. The method ofclaim 14, comprising controlling the amount of the bearing current usedto levitate the impeller transverse to the impeller axis of rotation soas to substantially minimize power consumption of the blood pump. 20.The method of claim 14, wherein the impeller impels the bloodcentrifugally and the blood pumped by the blood pump is output in adirection transverse to the impeller axis of rotation.
 21. The method ofclaim 14, comprising processing, with the controller, output from aplurality of hall sensors indicative of magnetic flux levels used tolevitate the impeller within the blood flow channel to determineeccentricity of the impeller within the blood flow channel, and whereinthe at least one impeller transverse position parameter is indicative ofthe determined eccentricity.
 22. The method of claim 14, comprisingestimating, with the controller. a pressure differential across theimpeller based on the at least one impeller transverse positionparameter.